The present invention relates generally to diagnostic imaging and, more particularly, to a method and apparatus of cardiac CT imaging using multi-spot emission sources.
Typically, in computed tomography (CT) imaging systems, an x-ray source emits a cone-shaped beam toward a subject or object, such as a patient or a piece of luggage. Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged. The beam, after being attenuated by the subject, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the subject. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis which ultimately produces an image.
Generally, the x-ray source and the detector array are rotated about the gantry within an imaging plane and around the subject. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point. X-ray detectors typically include a scintillator for converting x-rays to light energy adjacent the collimator, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom, although other types of detectors, including direct-conversion detectors, are known.
Typically, each scintillator of a scintillator array converts x-rays to light energy. Each scintillator discharges light energy to a photodiode adjacent thereto. Each photodiode detects the light energy and generates a corresponding electrical signal. The outputs of the photodiodes are then transmitted to the data processing system for image reconstruction.
It is generally desirable to have increased speed, coverage, and resolution of CT scanners, for example to improve imaging of the cardiac region. In recent years, manufacturers have improved scanners by increasing the gantry speed, by reducing the pixel size, and by extending the coverage of the detectors in the Z direction by extending the length of the detector in the Z direction. This approach has resulted in development of CT systems that have larger detectors. Detectors, in principle, may be extended in the Z direction to cover the entire cardiac region. However, such a length may be undesirable for a number of reasons. For instance, large detectors add cost and complexity to a CT system, not only in the detector components themselves, but in the data acquisition systems required to read out the increased number of channels. The increased detector size also results in an increased mass of the detector, thereby resulting in increased mechanical stresses in the components of the CT system.
As detectors get longer in the axial (Z) direction, an increase in the cone angle occurs as well. The cone angle is the angle, along the Z direction, between the focal spot and the edges of the detector. The increase in cone beam angle leads to cone beam artifacts in reconstructed images. Beyond a certain limit, the cone beam becomes severe, and increased scan coverage may not be accomplished by simply increasing the length of the detector along the Z direction.
A complete dataset is typically acquired during a rotation of a CT gantry through approximately 180 degrees, thereby defining the temporal resolution of a CT scanner, ignoring cone angles. Accordingly, the temporal resolution may be improved by spinning the gantry faster. However, mechanical stresses therein substantially increase with increased gantry speed, thereby imposing practical limits on the upper speed of the gantry.
Therefore, it would be desirable to design a CT apparatus and method to improve image quality of the cardiac region while increasing Z coverage of a subject.